Light-induced fluorescence has been identified as a powerful noninvasive method for tissue pathology recognition and monitoring [1]. (Numbers in brackets refer to references listed under REFERENCES below.) A major advantage of fluorescence imaging is that fluorophores can be detected at low concentrations using harmless, non-ionizing radiation that can be applied repeatedly to the patient.
In fluorescence imaging, the energy from an external light source is absorbed and almost immediately re-emitted at a longer, lower-energy wavelength. This energy is proportional to the electronic transition from the excited state to the ground state of the fluorophore molecule. Fluorescence that originates from chromophores naturally present in the tissue (endogenous) is known as auto-fluorescence. Fluorescence emission generally depends on the fluorophore's concentration, spatial distribution throughout the tissue, local microenvironment, and light attenuation due to differences in the amount of non-fluorescing chromophores.
Endogenous fluorophores are generally associated with the structural matrix of tissue (e.g., collagen and elastin), with cellular metabolic pathways (e.g., NAD and NADH), and with proteins associated with amino acids (e.g., tryptophan, tyrosin and phenylalanine). Cells in various disease states often undergo different rates of metabolism or have different structures associated with distinct fluorescent emission spectra. Table 1 shows a list of some of the tissue fluorophores with their respective peak excitation and emission wavelengths, as well as their common applications in studying various pathological parameters [2].
TABLE 1ExcitationPhysiologicMolecular sourcemax (nm)Emission max (nm)parameterTryptophan295345Proliferation,agingPepsin digestible335380Collagen cross-collagen cross-linkslinks, aging,glycationCollagenase370420Collagen cross-digestible collagenlinks, glycationcross-linksElastin cross-links400500Elastin cross-linksTyrosine275340InfammatoryinfiltrateCoproporphyrin405620P. acnesHorn365430Non-infammatoryacneNADH350450Mitochondrialactivity
Exogenous fluorophores have been developed as disease markers, for enhancing the pathological information in fluorescence images, and for photodynamic therapy applications. The fluorescing agent is either injected into the subject's body, consumed orally or topically applied over the region of interest. A variety of endogenous reporter probes have been used for enhanced detection of early cancers, including somatostatin receptor targeted probes; folate receptor-targeted agents; tumor cell-targeted agents; agents that incorporate into areas of calcification, bone formation or both; agents activated by tumor associated proteases; and topical photodynamic therapy [3-9].
Fluorescence measurements are usually carried out in the emission mode, i.e., radiation of a single wavelength is shone onto the skin and the emitted radiation (at wavelengths longer than the excitation wavelength) is measured and can then be analyzed. The most common technique to record fluorescence within a large tissue volume is associated with illuminating the tissue with a monochromatic or narrow spectral band plane wave (expanded light beam), and then collecting fluorescence signals emitted towards a CCD camera or other types of optical detectors [6, 10]. These methods can be generally referred to as planar methods and can be applied in epi-illumination or trans-illumination mode. Planar imaging has the added advantage that the same instrumentation can be used to image fluorescence in solutions and/or excised tissues. Planar fluorescence imaging is a useful technique when probing superficial structures (<5 mm deep), for example, during endoscopy, dermatological imaging, intra-operative imaging, tissue auto-fluorescence probing, or small animal imaging [11-15], with very high throughputs (speed of imaging).
Optical properties of skin have been studied and measured by numerous investigators [16-19] using spectroscopic or diffuse reflectance techniques. Values of the absorption coefficient, scattering coefficient, and anisotropy factor for various skin layers have been published. Table 2 lists the optical properties of various skin layers and corresponding sources of data.
TABLE 2Skin LayerOptical PropertyData SourceStratumScattering CoefficientGemert, Jacques, Sterenborgand Star [18]CorneumAbsorption CoefficientGemert, Jacques, Sterenborgand Star [18]Anisotropy FactorBruls and J. van der Leun [20]EpidermisScattering CoefficientGemert, Jacques, Sterenborgand Star [18]Absorption CoefficientJcaques [21]Anisotropy FactorBruls and J. van der Leun [20]DermisScattering CoefficientGemert, Jacques, Sterenborgand Star [18]Absorption CoefficientCui and Ostrander and B. Lee [22]Anisotropy FactorJacques, Alter and Prahl [23]
The absorption spectra of various chromophores have also been published by various researchers. Typical absorption spectra of major visible-absorbing pigments of human skin are shown in FIG. 1A. In the visible range, the main chromophores (light absorbers) of human skin are hemoglobin and melanin. Hemoglobin is found within the vascular network in the dermis while melanin is located in the epidermis. In certain skin types, like the East Asian skin type, absorption due to bilirubin is also observed. In the near-infrared (NIR) spectral range, these chromophores present very low absorption. The main light absorber in this spectral region is water.
FIG. 1B shows typical excitation/absorption spectra of the prominent endogenous fluorophores of skin, namely, collagen, elastin, flavins, NADH, porphyrins, tryptophan, pyridoxine, and lipo-pigments. FIG. 1C shows the emission spectra of the aforementioned skin endogenous fluorophores.
It should be noted that the term “porphyrin” as used herein may include porphyrin and/or its related compounds. Porphyrins are organic ring-shaped structures required in the synthesis of heme, a molecule necessary in the formation of hemoglobin. Numerous steps are involved in the pathways that result in porphyrin ring formation, and each step is dependent upon specific enzymes. Porphyrins are formed from coproporphyrin, which in turn forms pre-copropophyrin. The P. acnes bacteria seems to trigger this process and all three of them (porphyrins, coproporphyrin, pre-copropophyrin) can fluoresce, deriving their basic fluorescence capability from the porphyrin ring structure.
Diffuse optical tomography (DOT) is a means of deep-tissue optical imaging and has emerged as an important means of obtaining physiologically relevant tissue parameters such as blood oxygen saturation and total hemoglobin concentration [24]. DOT has a broad range of applications in optical breast imaging, functional brain spectroscopy and imaging, exercise medicine and photodynamic therapy monitoring [25-29]. In a typical DOT system, light is injected into tissue and is then detected at other points on the tissue surface. Three different measurement schemes are used for DOT imaging: time domain, frequency domain, and continuous wave (CW) [30-31]. Measurements are made either in transmission mode, reflection mode, or both. Of these three measurement types, the CW method is the simplest, least expensive, and provides the fastest data collection.
Fluorescence photography has been used to enhance the distribution of pigmentation, especially solar lentigenes, Propioni bacterium acnes, and open comedones [2, 32-35]. UV-excited fluorescence photography involves illuminating (excitation mode) the skin using a narrow-band UVA light centered at about 365 nm. The fluorescence image (emission mode) is captured using a color digital camera [2], where the camera lens is filtered so that it accepts only visible radiation (˜400-700 nm). The fluorescence image obtained includes the fluorescence of coproporphyrin produced by the bacteria with a maximum emission at 620 nm (appearing red) and the fluorescence of the lesions with impacted “horn” (contained in the comedones), which fluoresce throughout the visible spectrum with a white appearance. (Note that while it is unknown which molecule or molecules within the mixture known as “horns” fluoresce, any reference herein to the fluorescence of horns is understood to include the fluorescence of any such molecule found in horns.) Both of these sources of fluorescence appear on a background of diffuse blue-green fluorescence due to the emission by the dermal collagen cross-links [2]. A similar technique is also described in the following United States Patents and Published Patent Applications: 6961517, 6922523, 6317624, 20040146290, 20050195316, 20050201935, 20050203355, 20060092315, and 20060268148. A handheld device for determining skin age, proliferation status and photo damage level based on multi-spectral measurements is described in U.S. Published Patent Application 20070004972.
In UV-excited fluorescence photography, the channels of the UVA-excited color image, (i.e., red, green and blue) can be used and analyzed individually. For example, the blue channel can be used to detect hyper-pigmented spots that cannot be seen under normal white light imaging methods. The green and red channels have been used for acne detection and monitoring.
A method of assessment for inflammatory lesions using polarized light photography, with the polarizing filters oriented perpendicularly to each other, has been demonstrated [36, 37]. Cross-polarized and parallel-polarized imaging have been used for separating surface reflectance and diffuse reflectance, or in other words, for viewing surface and subsurface information individually. Comparing the image obtained using polarized light with the image obtained using UV-excited fluorescence photography has been used for identifying inflammatory lesions.
Known UV-excited fluorescence photography techniques, however, have several drawbacks. A significant drawback of the planar imaging technique is that it cannot resolve depth and does not account for non-linear dependencies of the signal detected on propagation depth and the surrounding tissue.
Additionally, it is wrongly assumed that the camera provides the required spectral separation between the red, green and blue color bands without any overlap. The response functions of CCD imagers, which are designed based upon the human visual response to colors, exhibit massive overlap among the three color channels. It is nearly impossible to separate the spectral information from the RGB image, even by using a very accurate camera model. Without such separation, accuracy is compromised when attempting to analyze the blue, green, and red channels for collagen, horns and porphyrin fluorescence, respectively.
Furthermore, many commercially available cameras use a Bayer pattern that is designed to make the captured image look aesthetically pleasing to the viewer. In doing so, scientifically significant spectral information within the data is corrupted.
Moreover, different fluorophores have different fluorescence yields, i.e., the efficiency of producing fluorescence. Simultaneously capturing the fluorescence signal from various fluorophores will hide the information of fluorophores having lower fluorescence yields.
Furthermore, superficial fluorescence activity may reduce the contrast of underlying activity from being detected owing to the simple projection viewing.
Tissue is a heterogeneous medium and the non-uniform distribution of chromophores therein will affect the fluorescence measurement of deeper fluorophores. The fluorescence signal captured by the camera, will be influenced by absorption due to the surrounding chromophores, mainly melanin and hemoglobin.
An additional limitation of known UV-excited fluorescence photography techniques is that they do not account for the non-uniformity of the illuminating light distribution in the fluorescence measurement and subsequent analysis. The areas of a tissue that receive more illumination may produce more fluorescence from a particular fluorophore compared to the lesser illuminated areas.
Considering the fluorescence from a particular fluorophore of interest as the signal and the fluorescence from other fluorophores as the noise, the above-discussed problems of known UV-excited fluorescence photography techniques lead to poor signal-to-noise ratios in the individual R, G, and B channels of images captured using such techniques. The simultaneous capture of various fluorescence spectral bands with overlapping detector response introduces noise.
Finally, the excitation (illumination) wavelength band selected is not optimized for imaging fluorophores of interest, such as collagen, horn, and porphyrin.
There is a need, therefore, for improved techniques that address the above-discussed drawbacks of the present UV-excited fluorescence photography state of the art.
An image capture apparatus described in U.S. patent application Ser. No. 11/167,540 to Budris et al., entitled “The Programmable Multi-spectral, Image Capture Environment,” and incorporated herein by reference in its entirety, provides a useful platform to target selected fluorophores and chromophores.
Patwardhan et al. have demonstrated that quantitative 3D reconstruction maps of tissue absorption and scattering can be obtained from time domain measurements [30, 38]. Using tissue simulating phantoms, Bjeorn, Patwardhan, et al., have also demonstrated the adverse effects of light absorption on fluorescence measurements and how the 3D optical property maps can be used to improve the fluorescence image quality [39].
3D reconstruction of optical properties, in particular oxy- and deoxy-hemoglobin, from multi-spectral measurements has been demonstrated on several occasions for applications such as breast cancer imaging, functional mapping of brain activity, and skin lesion imaging for melanoma [40-42].
International Patent Application PCT/US2007/063191 (WO/2007/103795) describes the use of multi-spectral imaging for demonstrating pre-emergent pimples using oxy-hemoglobin as an indicator. Various other applications of multi-spectral and/or hyper-spectral imaging are noted in U.S. Published Patent Application 20050270528. U.S. Pat. No. 6,208,749 describes a multi-spectral imaging approach for characterization of skin lesions. Further, the selection of wavelengths for multi-spectral absorption measurement can be optimized based on mathematical modeling of light/tissue interaction as demonstrated by Patwardhan et al, 2005 for multi-spectral skin lesion imaging applications [43]. 2D distribution maps of the chromophores melanin and hemoglobin have been obtained from the cross-polarized RGB images of skin using Independent Component Analysis (ICA), Principle Component Analysis (PCA), and color space transformations. 2D distribution maps of chromophores have also been obtained from multi-spectral measurements along with a mathematical model of light/tissue interaction [44-49]. Further images captured using UV and/or blue light illumination have been used to obtain distribution maps of melanin, while images captured using green light for illumination have been used to obtain distribution maps of hemoglobin.